JPH0654956B2 - Digital radiographic equipment - Google Patents

Description

TECHNICAL FIELD OF THE INVENTION The present invention relates to a digital radiography apparatus for irradiating a subject with radiation and converting the transmitted radiation dose into a digital signal for display.

[Technical Background of the Invention] FIG. 5 shows an image intensifier (hereinafter referred to as I.I.
Called. ) And an image pickup tube are configuration diagrams of a conventional digital radiography device. That is, X
The X-ray control unit 21 controls the X-ray tube 1 to irradiate the subject with X-rays. The X-rays that have passed through the subject P become a radiation image I.I. I. It is incident on 22. I. I. 22 amplifies this radiation image and outputs it as an optical image. This optical image is input to the image pickup tube 24 such as a bicicon via the optical system 23. The image pickup tube 24 converts this optical image into an analog video signal and supplies it to the A / D converter. A / D converter 7
Converts the analog video signal into a digital video signal and supplies it to the image memory 28. The image memory 28 temporarily stores this digital video signal. In the image processing unit 26, the digital video signal stored in the image memory is
Digital image processing such as digital subtraction is performed. This digital image-processed digital signal is D
It is supplied to the / A converter 10. The D / A converter 10 converts this digital signal into an analog signal and displays it on the TV monitor 29.

I. I. The conventional digital radiography apparatus using the image pickup tube and the image pickup tube has the following problems.

That is, I. I. Is a vacuum tube with a built-in electron lens, so there is a limit to increasing the visual field due to its structure.
For example, a commonly used one is 12 inches (30
cm) up to about φ, the field of view is circular.

I. I. Since the input surface is a vacuum tube having a convex or concave shape, distortion such as thread winding distortion occurs in the output image. Further, even if the uniform absorbing object is photographed, the density distribution of the output image is not uniform, and so-called "sewing" occurs especially in the peripheral portion of the image.

The image pickup tube has a narrow dynamic range and cannot be used for both fluoroscopy and photography.

SUMMARY OF THE INVENTION The present invention has been made to solve the above problems, and provides a digital radiography device having a wide dynamic range, being lightweight and compact, capable of ensuring a large field of view, and causing no distortion in an output image. The purpose is to do.

[Summary of the Invention] In order to achieve this object, the present invention provides (1) a radiation source for irradiating a subject with radiation and a radiation transmitted through the subject, and generates an electric charge according to the intensity thereof. A one-dimensional or two-dimensionally arranged radiation-electric conversion unit, a first capacitor connected in parallel with each radiation-electric conversion unit, a second capacitor connected in parallel with the first capacitor, and a first capacitor A connection control unit for controlling connection between the capacitor and the second capacitor, a power supply connected to one terminal of each first capacitor, a switch unit provided at each other terminal of each first capacitor, and sequentially closed, and a switch unit An output terminal that outputs a current flowing from the power source to the capacitor when the capacitor is closed, an A / D conversion unit that converts the amount of current generated at the output terminal into a digital signal, and the digital signal A digital radiography device (2), characterized in that it comprises a display section as shown, wherein the output of the radiation source is switched to increase in connection with the connection of the second capacitor. (3) The digital radiography device according to claim 1, wherein the capacity of the second capacitor is larger than the capacity of the first capacitor. (4) The digital radiography device according to claim 1 (4) The digital radiography apparatus according to claim 1, wherein a diode and a Wener diode are connected in parallel.

[Embodiment of the Invention] An embodiment of the present invention will be described below with reference to the drawings. FIG. 1 is a block diagram of the same embodiment. In the figure, the high voltage generator 12 generates electric power for driving the X-ray tube 1 according to a control signal from the controller 13.

The power supplied from the high voltage generator 12 to the X-ray tube 1 is roughly divided into two: power for irradiating weak X-rays for fluoroscopy and power for irradiating strong X-rays for imaging. There are types. And the power of see-through is 100 compared with the power of shooting.
It is less than a fraction. The X-rays emitted from the X-ray tube 1 pass through the subject 2 and enter the imaging plate 3 such as an intensifying screen that converts the X-ray energy into light energy. The X-ray dose incident on the imaging plate 3 corresponds to the remaining amount that could not be absorbed by the subject 2, and this becomes an X-ray image. This incident X-ray is converted into an analog electric signal according to the X-ray dose by a photoelectric conversion element formed of a two-dimensional array in the detection unit 4. The electrically converted analog signal is digitally converted by the A / D converter 7 in time series and stored in the image memory 8. The image memory 8 can store data for one image or several images.
Data is sequentially stored at a specific address by a control signal from. The arithmetic processing unit 9 retrieves data from the image memory 8 and arithmetically operates it, and returns the result to the image memory 8 again. The calculated data of the image memory 8 is the D / A converter 10
Is converted into an analog signal. This analog signal is displayed on the image monitor 11 as an X-ray image.

Next, the detection unit 4 that converts the amount of light emitted by X-rays into an electric signal
Will be described in more detail with reference to the block diagram of FIG.

Element e1, including a photoelectric conversion element that converts light into electricity
1 to e128,128 are evenly arranged in a two-dimensional array. 128 elements in a horizontal row of the elements e1,1 to e128,128 are commonly connected by a single common signal line via a switch. The same 128 horizontal elements are arranged in the vertical direction as well. That is, there are 128 common signal lines. Each of the common signal lines is connected to the input of the multiplexer 6.

The multiplexer 6 sequentially selects one of the 128 common signal lines according to the control signal CK ₂ from the controller 13 and outputs it to the A / D converter 7 as an output signal of the detector 4.

In this way, the detection unit outputs 128 common signals in time series from the multiplexer 6. Actually 128 x 12
8 and 2048 × 2048 or the like are used, and the output of the multiplexer 6 is also plural. However, for simplicity, the above example will be described.

Next, FIG. 3 shows a circuit diagram of a part of the horizontal 128 elements of the detection unit 4. The positive electrodes of the photoelectric conversion elements Pd1 to Pd128 are connected to the negative electrode of the power source E. Also, capacitors Cd1 to C
The d128 is connected in parallel to the photoelectric conversion elements Pd1 to Pd128. Further, capacitors Ce1 to Ce128 having a larger capacity than the capacitors Cd1 to Cd128 are connected in parallel with the capacitors Cd1 to Cd128 via the switches c1 to Sc128. Switch Se1 connected in series with capacitors Ce1 to Ce128
˜Se128 are opened / closed by a control signal CK3 from the control unit 13. For fluoroscopy with a small X-ray dose, the switches Sc1 to Sc128 are opened. Further, when the X-ray dose is large, the switches Sc1 to Sc128 are closed and the capacitance of the capacitors connected in parallel to the photoelectric conversion elements Pd1 to Pd128 is increased. Thus, the capacitors Sc1 to Sc128
If the X-rays are connected, the capacitance will be large, so that the charged charges will not be saturated even when the incident X-ray dose is large, and even if the incident X-ray dose is small, opening the switches Sc1 to Sc128 will damage the S / N ratio. I don't care. Capacitors Cd1 to Cd1
The capacitance Cd of 28 is the capacitance of a capacitor suitable for transparent X-rays. In addition, the capacitance of the condenser suitable for X-rays for photography is C0.
Then, the capacitance Ce of the capacitors Ce1 to Ce128 is an effective capacitance including the stray capacitance of the switch, and Ce = C0-Cd.

The cathodes of the photoelectric conversion elements Pd1 to Pd128 are each preamplifier A1.
.. to A128, the switches Sd1 to Sd128 on the photoelectric conversion element side are connected at the same time. Then, when this switch is closed, the electric charge of the power supply voltage E is charged in the capacitor. The preamplifiers A1 to A128 are respectively Pd1 to
The cathode potential of Pd128 is input with high impedance, amplified, and then output with low impedance. Preamplifier A1
The outputs of A to A128 are connected to the switches Sa1 to Sa128 on the output side. The shift register F receives the control signal CK1 from the control unit 13 and switches Sd1 to Sd128 and Sa1 to
It controls the opening and closing of Sa128. A signal corresponding to the charges accumulated by the control of the shift register F is output to the common signal line in time series.

Next, the operation of this embodiment will be described. First, in the case of see-through, the switches Sc1 to Sc128 are in the open state. The capacitors Cd1 to Cd128 are supplied with electric charge by the power supply voltage E and charged to the inter-terminal voltage. The X-ray tube 1 irradiates the subject with weak continuous X-rays having a tube current of several mA. The X-rays that have been attenuated and transmitted by the subject are converted into light by the imaging plate 3 and are incident on the diodes Pd1 to Pd128, respectively, to generate electric charges according to their respective intensities.
These charges flow into the capacitors Cd1 to Cd128 and cancel the precharged charges. That is, the charge proportional to the incident X-ray intensity decreases. During transmission, switch Sa1
~ Sa128 are sequentially closed by the shift register F synchronously. When the switch Sa1 is closed, a current is supplied from the power supply voltage E to the capacitor Cd1 by the amount of the charge canceled by the X-ray. This current is detected at the output terminal OUT and output to the A / D conversion unit 7. While the switch Sa1 is closed, the capacitor Cd1 is charged to the initial state. Next, the switch Sa2 is closed, and the X-ray dose incident on the diode Pd2 is similarly detected. In this way, while the X-ray is being irradiated, the diodes Pd1 to Pd1
The X-ray dose incident on 28 is sequentially and synchronously detected.

All switches Sc1 to Sc128 are closed at the time of photographing.
Then, as in the case of the perspective, electric charges are supplied to the capacitors Cd1 to Cd128 and the capacitors Ce1 to Ce128 by the power supply voltage E to be charged to the inter-terminal voltage. The X-ray tube 1 bombards the subject 2 with a strong X-ray pulse having a tube current of several hundred mA. The X-ray transmitted through the subject 2 is diode Pd1 as in the case of fluoroscopy.
~ Pd128 converted to current. These currents cancel the charge stored in the capacitors Cd1 to Cd128 and Ce1 to Ce128. Unlike the case of the fluoroscopy, the switches Sa1 to Sa128 are sequentially closed only once for each X-ray pulse. Then, when each of the switches Sa1 to Sa128 is closed, an output corresponding to the intensity of the incident X-ray is obtained at the output end OUT as in the case of fluoroscopy.

The time during which the switches Sa1 to Sa128 are closed requires a length of time until the capacitors are charged to the initial state. Therefore, the capacitance of the capacitor is small, and a large amount of charge cannot be stored. In the case of fluoroscopy, the time the switch is closed is shorter than in the case of shooting.

Next, a second embodiment of the detector 4 and the circuit diagram of FIG. 4 will be described. The difference between the first embodiment and the second embodiment is that a diode and a Zener diode are used instead of the switches Sc1 to Sc128 to automatically switch the capacitance. Diodes D1 to D128 and Zener diodes Z1 to Z128
Are connected in parallel. Capacitors Ce1 to Ce128 are connected in series with these. Those connected in series are connected in parallel to the capacitors Cd1 to Cd128. Further, the capacitors Cd1 to Cd128 are photoelectric conversion elements Pd1.
~ Pd128 connected in parallel. The reading operation in the accumulation mode shown in FIG. 2 is performed as follows. Once switch Sd1 ~ S
d128 and Sa1 to Sa128 are closed, and capacitors Cd1 to Cd
After charging d128, open the switch. When the power supply voltage E is 20 (V), the voltage between the terminals of the capacitors at this time is 20 between the terminals of the capacitors Cd1 to Cd128.
(V), the voltage between the terminals of the capacitors Ce1 to Ce128 is 19. V because of the voltage drop (0.7 V) of the diodes D1 to D128.
It becomes 3 (V). The electric charge canceled by the incident X-ray starts from the electric charge of 20 (V) charged in the capacitors Cd1 to Cd128. Assuming that the Zener voltages of the Zener diodes Z1 to Z128 are 12 (V), the Zener diodes Z1 to Z1 until the voltage of the capacitors Cd1 to Cd128 falls to 7.3 (V).
Since Z128 does not conduct, the voltage of the capacitors Ce1 to Ce128 does not drop. However, when the voltage of the capacitors Cd1 to Cd128 becomes 7.3 (V) or less, the voltage between the terminals of the Zener diode becomes higher than the Zener voltage, the Zener diodes Z1 to Z128 become conductive, and from the capacitors Ce1 to Ce128 to the capacitors Cd1 to Cd128. A current flows and the capacitor Ce1〜
The voltage of Ce128 goes down. As described above, in the second embodiment, the capacitors Ce1 to Ce128 are automatically used when the amount of the canceled charges is increased. That is, the capacitance can be automatically switched by using the Zener diodes Z1 to Z128 and the diodes D1 to D128.

[Effects of the Invention] As described above, according to the present invention, at least two capacitors are connected in parallel to a photoelectric conversion element, and the storage capacity is changed according to the irradiation dose of X-rays, thereby supporting a wide range of X-ray doses. It is possible to provide a digital radiography device having a dynamic range that can be obtained.

[Brief description of drawings]

FIG. 1 is a block diagram of an embodiment of a digital radiography apparatus according to the present invention, FIG. 2 is a block diagram of a detection unit of the same embodiment, and FIG. 3 is a circuit of a first embodiment of elements of the detection unit. FIG. 4 and FIG. 4 are circuit diagrams of a second embodiment of the elements of the detection section, and FIG. 5 is an overall configuration diagram of a conventional digital radiography apparatus. 1 ... X-ray tube, 4 ... Detector Pd1 to Pd128 ... Photoelectric exchange element, Cd1 to Cd128 ... Capacitor C1 to Ce128 ... Capacitor, Sd1 to Sd128 ... Switch Sc1 to Sc128 ... Switch, A1 to A128 ...... Preamplifier D1 to D128 …… Diode, Z1 to Z128 …… Zener diode

Claims (4)

[Claims] 1. A radiation source for irradiating a subject with radiation, and a one-dimensional or two-dimensional array of radiation-electric converters for detecting the radiation transmitted through the subject and generating electric charges according to the intensity thereof. A first capacitor connected in parallel with each of the radiation-electric conversion units, a second capacitor connected in parallel with the first capacitor, and a connection control unit controlling connection between the first capacitor and the second capacitor And each first
A power source connected to one terminal of the capacitor, a switch section provided at each other terminal of each first capacitor and sequentially closed, and an output for outputting a current flowing from the power source to the capacitor when the switch section is closed. A digital radiography apparatus comprising: an end, an A / D conversion unit that converts the amount of current generated at the output end into a digital signal, and a display unit that displays the digital signal. 2. A digital radiography device according to claim 1, characterized in that the output of the radiation source is switched to increase in connection with the connection of the second capacitor. 3. The digital radiography apparatus according to claim 1, wherein the capacity of the second capacitor is larger than the capacity of the first capacitor. 4. The digital radiography apparatus according to claim 1, wherein the connection control unit connects a diode and a Zener diode in parallel.